Method and apparatus for quantitative bone matrix imaging by magnetic resonance imaging

ABSTRACT

An apparatus and technique for measuring volumetric, 3-D bone organic matrix density is described. The techniques includes providing a first pulse sequence fragment selected to suppress at least two fluid resonance signals and providing a second pulse sequence fragment which images at least solid signals. In one embodiment, a series of RF pulses used to suppress fluid in the bone marrow spaces, particularly in cancellous bone tissue, is combined with solid state projection reconstruction magnetic resonance imaging (MRI) to provide a three-dimensional image of a bone in which the dominant signal arises substantially from solid organic bone matrix.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a National Stage application, of and claims thebenefit of PCT Application No. PCT/US03/15801, filed on May 19, 2003,which claims the benefit of U.S. Provisional Application No. 60/381,161,filed on May 17, 2002, which are incorporated herein by reference.

FIELD AND BACKGROUND OF THE INVENTION

This invention relates generally to magnetic resonance imaging (MRI) andmore particularly to a method and apparatus for measuring the degree ofvolumetric (3-D) bone organic matrix density.

Glossary of Currently Used Terms to Describe the Structure, Compositionand Organization of “Bone”

The term “bone” may be used to refer to various levels of bonestructure, composition and organization, from the gross visual, nakedeye identification of a specific whole bone or a part of a whole bonesuch as the femur (the upper bone of the thigh), viz., “bone” as anorgan or a part of an organ, to the specific structural organization ofportions of a bone by light microscopy, that is, of “bone” as a tissue,e.g., compact bone or cancellous (spongy or trabecular) bone, or even tothe organization of the individual components of bone tissue, that is,“bone” as a substance or material (bone substance) whose individualcomponents can be visualized, for example, by electron microscopy andother techniques.

FIGS. 1A, 1B and 1C show the cylindrical portion of the diaphysis(shaft) of an animal's tibial bone (portion of a whole gross bone). Thebone material or substance is densely packed—thus the cylindrical wallof the bony shaft, the bone tissue, is described as compact bone. At theends of a long bone, the bone as an organ, for example, the head andneck regions of the femur, the bone substance is not densely packed.Indeed, the bone tissue is clearly organized in a specificthree-dimensional, loosely packed interconnected open network consistingof small segments of bone substance referred to as trabeculae, and thebone tissue referred to as cancellous or spongy or trabecular bonetissue.

The basic bone substance or material is composed principally of a softmatrix consisting primarily of the (fibrous) protein collagen and smallamounts of other organic constituents and other extracellular,extravascular organic constituents and is referred to as the organicmatrix of bone substance. It is into this matrix that the other majorcomponent of bone substance, the hard crystals of calcium-phosphate(“apatite”) (solid mineral phase), is deposited. Therefore, the twomajor components of bone substance are: (1) a soft organic matrix and(2) the hard solid phase of the calcium-phosphate crystals, the solidmineral phase. These two components provide a good portion of themechanical properties of bone as an organ, a tissue and a materialsubstance, as well as many of the physiological functions of bonesubstance. The organic matrix of bone substance is ordinarily consideredto consist of the extracellular, extravascular organic components of thebone substance and the bone tissue, and is chemically analyzed bymeasuring the collagen or collagen and other known proteins and organicconstituents of the extracellular, extravascular matrix.

The composition of the bone substance can therefore be expressed interms of the relative proportions of the two major constituents of bonesubstance, either by the weight percentages of the two major components,viz., the weight percentage of the organic matrix compared with theweight percentage of the solid mineral phase, or better still, byconsideration of the volumetric density of the solid mineral phase and,independently, of the volumetric organic matrix density: it is much moreimportant and informative to be able to express the composition of bonesubstance and of bone tissue in terms of volume, e.g., the mass orweight of the solid mineral phase in a unit volume of bone substance(gm/cm³), and the volumetric density of the mass or weight of theorganic matrix in a unit volume of bone substance in gm/cm³. From thesedata, viz., the volumetric density of the solid calcium-phosphatemineral phase and the volumetric density of the soft organic matrix, itis possible to calculate the extent or degree of mineralization, thatis, the extent to which a unit volume of bone substance in compact orcancellous bone tissue is mineralized. This has not been accomplishedpreviously by any noninvasive technique but has been accomplished asdescribed in this patent application. Because this inventionaccomplishes the measurement of bone organic matrix density by MRI, noionizing radiation is used, and therefore patients may be examinedrepeatedly without the risk associated with ionizing radiation.

We therefore use the following terms and definitions: (1) bone mineraldensity; (2) bone organic matrix density (organic matrix density ofbone); (3) extent (or degree) of bone mineralization. It is important todistinguish, however, whether these data have been calculated from 2-D(areal) measurements and not measured as 3-D volumetric data, or havebeen directly measured as volumetric 3-D data, as accomplished by thetechniques described in this patent application. It is from thesedirectly measured volumetric data that it is possible to calculate thevolumetric extent or degree of mineralization of bone substancenon-invasively. The ratio of volumetric, 3-D bone mineral density tovolumetric bone organic matrix density is referred to as the extent ordegree of bone mineralization.

Currently, the 3-D volumetric extent or degree of mineralization cannotbe measured noninvasively. It can be determined by chemical andgravimetric analyses of a piece of bone tissue removed from a patient oran animal surgically by biopsy. Clearly, such a surgical procedurecannot be carried out each time a measurement is needed to follow thecourse of any particular disease or to assess the efficacy of aparticular drug or treatment over a long, extended period of time.

Techniques for measuring bone mineral density non-invasively have beendeveloped. Two such techniques are by X-ray and magnetic resonanceimaging (MRI). Prior to this invention, no noninvasive methods formeasuring volumetric bone matrix density had been described.

Currently, two of the most commonly used techniques to measure bonemineral density are: (1) dual energy x-ray absorptiometry (DXA) and (2)computed tomography (CT). DXA utilizes x-rays of two energies. Themineral and soft tissue each exhibit different x-ray scattering crosssections at each energy level, enabling a map of mineral density to becomputed from the scan data. However, because of the variablecomposition of the soft tissue and its variable depth along the viewdirection, overlapping bone structure and the inhomogeneity of the 3-Dspatial distribution of the trabeculae in cancellous bone, for example,the most commonly analyzed bone tissue using this technique, the 2-D(areal) measurement of bone mineral density may not reflect the truevolumetric 3-D bone mineral density. Indeed, serious questions have beenraised in the literature about the validity and usefulness of the dataobtained by DXA.

Computed tomography (CT) produces an accurate measurement of volumetric,3-D bone mineral density (grams per cubic centimeter). However, when thex-ray intensity is sufficient to make the CT scan quantitativelyaccurate (quantitative CT or QCT), the radiation dose to the patient ishigh, limiting the number of scans permissible for a single patient,thus preventing the use of QCT on women of child bearing age, growingchildren and patients who may require repeated measurements in order tofollow and assess the course of a disease or injury or to assess theefficacy of treatment. Like any x-ray based measurement, CT does notdistinguish bone matrix from soft tissue, and is susceptible to errorsbecause of the variability of soft tissue composition and depth.

Measurement of the Volumetric (3-D) Density of the Organic Matrix ofBone by Proton Nuclear Magnetic Resonance Imaging

Magnetic resonance imaging (MRI) is a widely used and highly effectivemeans of producing two and three dimensional images of the body. Withsuitable settings of the parameters of the pulse sequence (e.g. thedefinition of the timing, amplitude, frequency, phase and other detailsof the radio frequency and magnetic field gradient pulses and variouscontrol functions produced by the scanner) MRI can yield quantitativedata on certain properties of tissues.

Conventional proton MRI detects the fluid proton (hydrogen) content ofsoft tissues, which is mainly liquid water, and to a lesser extent, fat.However, most solid substances, including bone substance, do not yieldsignals in conventional MRI, and thus are not detected and therefore notvery visible in conventional MRI. Conventional MRI therefore yields noinformation on the composition of the bone substance To obtain MR imagesof solid materials, it is necessary to utilize specialized pulsesequences and scanner hardware.

The Larmor (resonance) frequency of a nuclear spin is proportional to aconstant (the magnetogyric ratio) specific to each nuclear species, aswell as proportional to the total magnetic field in which it isimmersed. During MR scanning nuclear spins in a material or bodyexperience both the strong magnetic field of the scanner as well as thesmaller local magnetic fields of neighboring nuclear spins. Theinstantaneous sum of all these fields at the site of a particular spindetermines its Larmor frequency. In the most general case of a solid, aparticular spin will experience a local field dependent on the number,spatial location and quantum state of all the other spins in itsvicinity, in addition to the much stronger field of the scanner. Thiseffect is called dipole-dipole, or spin-spin, coupling. Different spinswill experience somewhat different local fields, but all spins willexperience the same scanner field. Therefore the Larmor frequencies willbe distributed about a central value (determined by the scanner field),yielding a frequency spectrum having a finite line width.

If the spins are widely separated from each other in the material, thespectral line width will be relatively small. If the spins are in fastrandom relative motion with respect to each other (as in the case of afluid such as liquid water), such that the local field is rapidly timedependent with a short autocorrelation time, the effective local fieldis averaged to zero, and the line width will be very small. For the caseof typical organic solids, the proton line width may be on the order ofseveral thousand Hertz (Hz) to several tens of kilohertz (kHz). Forfluid systems the proton frequency spectrum line width may be on theorder of less than 1 Hz to a few hundred Hz.

The line width strongly affects the performance of MR imaging. Theinverse of the spectral line width is generally known as T₂ if thespectral broadening is due to spin-spin coupling among like(homogeneous, e.g., all proton) spins, and T₂* if the spectralbroadening is due to a static (not time dependent) distribution of theintrinsic Larmor frequency unrelated to homogeneous spin-spin coupling.Both types of broadening mechanisms are often present, and T₂* isusually used to denote the total broadening of the spectral line.Therefore, in the remainder of the invention description T₂* will beused in the conventional manner to encompass all the spectral linebroadening effects intrinsic to the subject or specimen as well as dueto scanner main magnet inhomogeneity. T₂* represents the characteristictime for the MR signal to dephase following an RF pulse, and imposes alimit on both the spatial resolution and the signal-to-noise ratioobtainable in an image. Because the spectral line widths of solidmaterials, such as bone matrix, are usually far larger than those offluid materials, such as tissue water or fat, it is expected that imagesof solid materials are of much lower spatial resolution andsignal-to-noise ratio.

Following the initial RF pulse which elicits a transverse magnetization(the detection and recording of which constitutes the MR signal), themagnetization begins to dephase under the dipole-dipole coupling andother mechanisms. Essentially all conventional MR pulse sequencesrequire the generation and recording of a spin echo, which is the forcedrephasing of the dephased signal by the application of a magnetic fieldgradient pulse reversal or a 180 degree (180°) RF pulse. The T₂* type ofbroadening may be overcome with the 180° pulse, but the T₂ type ofbroadening cannot be overcome by any simple sequence of RF pulses.Because of certain scanner engineering constraints, the minimum timebetween the initial RF pulse and the spin echo (the echo time, or TE)can usually be no shorter than a few milliseconds, which limits the linewidths to be no greater than a few hundred Hz if the signals are to bedetected. Therefore, a solid material (operationally defined here as anysubstance with T₂ less than about 1 ms), including but not limited tobone matrix, cannot be imaged with conventional MR pulse sequences.

One MRI technique which can be used to image solids is referred to assolid state MRI or projection reconstruction MR imaging or projectionMRI or even more simply projection imaging. Projection imaging employsonly an initial RF pulse and does not elicit spin echoes. The RF pulseelicits a magnetization response referred to as a free induction decay(FID). The FID begins to de-phase immediately following the end of theRF pulse. The FID is ignored in conventional pulse sequences, but may infact be recorded to create a data set that can be reconstructed into animage. In simple projection imaging (without slice selection), the FIDis recorded in the presence of a constant amplitude magnetic fieldgradient. The magnetic field gradient direction is advanced to a newdirection, and the FID is elicited with an RF pulse, and recorded again.The process is repeated to cover all directions in three dimensionalspace, and the image is reconstructed from the recorded data using oneof several possible algorithms to yield a three dimensional image.Because no spin echo is required, short T₂ or T₂* does not prevent theprojection method from imaging solid materials. However, short T₂ or T₂*will still reduce the spatial resolution and signal-to-noise ratio inprojection imaging just as they do in conventional MRI.

Although simple projection imaging can be used to make images of bonematrix, it has a limitation in that all proton containing substances,including bone marrow, will be imaged. Because the marrow proton signalslargely arise from fluid substances (e.g. water and fat), their signalswill be imaged at relatively high signal-to-noise ratio, and willdominate the signals from the matrix. Thus, since solid state MRImeasures both fluid and solid constituents, it cannot be used toeffectively measure the volumetric, 3-D density of the organic matrix ofbone.

In summary, conventional proton MRI detects only the fluid constituentsof bone and thus cannot be used to measure the volumetric, 3-D densityof the organic matrix of bone. Solid state MRI, on the other hand,detects both the solid and fluid constituents of a bone and since thefluid constituents obscure the solid constituents, solid state MRIcannot be used to measure the volumetric, 3-D density of the organicmatrix of bone. Thus, neither conventional proton MRI nor unmodifiedsolid state MRI can be used to measure the volumetric density of theorganic matrix of bone.

SUMMARY OF THE INVENTION

It has, in accordance with the present invention, been recognized thatit would be very advantageous to provide a method for measuring the 3-Dvolumetric density of the organic matrix of bone since this would makepossible to combine the data so obtained with the 3-D volumetric densityof the bone mineral to measure the volumetric 3-D extent or degree ofmineralization in bone substance.

In the present invention, projection imaging is combined with a fluid(e.g. water and fat) signal suppression technique to produce images ofonly the volumetric, 3-D bone organic matrix thereby enabling themeasurement of the volumetric, 3-D bone organic matrix density. Thewater and fat signal suppression technique is selected such that waterand fat signals are mostly eliminated, while the signals arising fromthe organic matrix are largely retained. In accordance with the presentinvention, the technique for measuring the volumetric, 3-D bone organicmatrix density includes providing a first pulse sequence fragmentselected to suppress at least two fluid resonance signals and providinga second pulse sequence fragment which images at least solid signals.

With this particular arrangement, a technique which enables quantitativethree-dimensional (3D) imaging of bone organic matrix density isprovided. By using a first pulse sequence fragment which suppress atleast two fluid resonance signals and then using a second pulse sequencefragment which images at least solid signals, the technique of thepresent invention can be used to non-invasively measure the volumetricdensity of the bone organic matrix (in grams per cubic centimeter) overan extended volume of a particular bone. The technique of the presentinvention does not rely on an invasive procedure (e.g. biopsy) nor doesit rely on any procedure which utilizes x-rays or other ionizingradiation. The technique of the present invention thus presentsrelatively little, if any, risk to the subject.

By combining a specific series of narrowband radio frequency (RF) pulsesto suppress fluid (e.g. water and fat) in bone marrow with proton solidstate projection reconstruction MRI, a three dimensional image of thebone in which the dominant signal arises only from the solid boneorganic matrix is provided. The narrowband RF pulses substantiallysuppress the signals from the water, fat and other relatively mobileconstituents within the marrow spaces and in the bone substance,especially in cancellous bone, while retaining a significant portion ofthe signals from the solid immobile organic matrix constituents of thebone substance. The present invention thus enables the quantitativethree-dimensional imaging of bone organic matrix density noninvasivelywith solid state MRI.

Water and fat suppressed projection MR imaging thus utilizes the largedifference between the proton T₂*s of the solid organic matrix and thefluid constituents of bone to suppress the fluid signals whilepreserving the solid organic matrix signals. The solid constituentsinclude collagen and other proteins and organic constituent signals,some molecularly immobile water, and exhibit very short T₂*. The fluidconstituents include molecularly mobile water and fat, with long T₂*. Inthe technique of the present invention, chemical shift selective lowpower 90 degree pulses excite mobile water and fat magnetization whichis subsequently dephased by gradient pulses, while the magnetization ofcollagen and other solid immobile organic matrix constituents of bonematrix and immobile water remains mostly in the z-direction. Additionalselective 180 degree pulses in alternate scans further cancel theresidual water and fat magnetization. Following water and fatsuppression and acquisition of a proton bone matrix density image, thistechnique can be used in combination with volumetric, 3-D bone mineraldensity measurement by solid state ³¹P projection MRI to determine thevolumetric, 3-D degree or extent of bone mineralization, that is, thefraction of the total volume of bone substance occupied by the bonemineral or the grams of bone mineral in a unit volume of bone substance.

Thus, by positioning a bone in a substantially static magnetic field andsubjecting the bone to a first pulse sequence fragment selected tosuppress at least two fluid proton MR signals and then subjecting thebone to a second pulse sequence fragment to acquire at least a solidstate proton MR image it is possible to acquire proton RF signalsemitted by the bone. The proton RF signals are processed to generatedata representative of the volumetric, 3-D bone organic matrix density.An additional pulse sequence may be used to acquire the solid state ³¹pimage of the bone. The 3-D volumetric bone mineral density and the 3-Dvolumetric bone organic matrix density data are then used to calculatethe degree or extent of volumetric bone mineralization.

In accordance with a further aspect of the present invention, a magneticresonance imaging (MRI) system for obtaining a quantitative bone organicmatrix density measure includes an MRI pulse sequence generator forproviding a first pulse sequence fragment selected to suppress at leasttwo fluid resonance signals and for providing a second pulse sequencefragment which images at least solid signals.

With this particular arrangement, a system which provides a quantitativethree dimensional image of bone organic matrix density is provided. Inone embodiment, the MRI pulse sequence generator combines a series ofnarrowband radio frequency (RF) pulses with proton solid stateprojection reconstruction MRI to provide a three dimensional image ofthe bone in which the dominant signal arises only from the solid organicbone matrix. The narrowband RF pulses substantially suppress signalsgenerated in response to fluids (e.g. water and fat) in bone marrow,while retaining a significant portion of the matrix signal.

Normally the solid state projection reconstruction MRI, especially ofcancellous bone, would image the marrow tissue and spaces (whichcontains a significant amount of fluid and fat). However, since thenarrowband RF pulses suppress a substantial amount of the fluid signals,the proton solid state projection reconstruction MRI provides athree-dimensional image of the bone substance in which the dominantsignal arises primarily, if not totally, from the solid organic bonematrix. The system can thus provide a quantitative three-dimensionalimage of organic bone matrix density non-invasively with protonsolid-state magnetic resonance imaging. The volumetric bone organicmatrix density can then be used with data from the volumetric bonemineral density to obtain a measure of the degree of bone mineralizationin a selected portion of either compact or cancellous bone tissue.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing features of the invention, as well as the invention itselfmay be more fully understood from the following detailed description ofthe drawings, in which:

FIG. 1A, is a large portion of a tibial bone of an animal and serve asan example of bone as an organ;

FIG. 1B is a cross sectional view of the bone shown in FIG. 1A and showsthe dense compaction of the bone substance of which the organ iscomposed which is referred to as com-pact bone tissue;

FIG. 1C is a section through a human femoral head (H), neck (N), andshaft (S); it should be noted that the femoral head, neck and portion ofthe trochanter consist of small pieces of bone substance (trabeculae)packed loosely in a sponge-like manner; this is referred to ascancellous bone and can be compared with the very dense compact bone inthe shaft of the bone;

FIG. 2 is a block diagram of a system for measuring bone matrix density;

FIG. 3 is a flow diagram showing the steps in a technique for measuringbone matrix density in bone substance; and

FIG. 4 is comprised of FIGS. 4A and 4B which are plots of radiofrequency (RF) and magnetic field gradient pulse sequences,respectively, used to measure bone organic matrix density;

FIGS. 5A and 5B are plots of transverse and longitudinal componentsrespectively of magnetization M following rectangular RF pulses;

FIG. 6A is a schematic illustration of a phantom of corn oil and water;

FIG. 6B is a schematic illustration of a single pulse proton spectrum ofthe phantom;

FIG. 6C is a schematic illustration of a spectrum of the phantomfollowing application of the water and fat suppression sequence;

FIG. 6D is a schematic illustration of a one dimensional total protonprojection image profile of the phantom without water and fatsuppression;

FIG. 6E is a schematic illustration of a water and fat suppressedprojection imaging profile of the phantom;

FIG. 6F is a single plane from a 3D total proton projection image of thephantom;

FIG. 6G is a single plane from a 3D water and fat suppressed projectionimage of the phantom;

FIG. 7A is a pair of single planes from a 3D total proton projectionimage of a bovine femur diaphysis with intramedullary (IM) fat intact inthe intramedullary cavity;

FIG. 7B is a pair of single planes from a water and fat suppressedprojection image of a bovine femur diaphysis with intramedullary (IM)fat intact in the intramedullary cavity showing that the signal from allmaterial in the intramedullary cavity is suppressed.

FIG. 7C is a photograph of the a bovine femur diaphysis withintramedullary (IM) fat;

FIG. 8 is a series of single planes from the proton total and 3D waterand fat suppressed projection images of dry tendon and trabecular bone.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

Before describing a system for quantitative bone matrix imaging by solidstate proton magnetic resonance imaging (MRI) system and the operationsperformed to produce an image which aids in the diagnosis and treatmentof osteoporosis and other conditions, some introductory concepts andterminology are explained.

As mentioned above, for the case of typical organic solids, the protonline width may be on the order of several thousand Hertz (Hz) to severaltens of kilohertz (kHz). For fluid systems, the proton line width may beon the order of less than 1 Hz to a few hundred Hz. Thus, the term“fluid,” as used herein refers to any material yielding nuclear magneticresonance (NMR) signals (having a spectral response) below 1 kilohertz(kHz) and the term “solid” as used herein refers to any material havinga spectral response at or above 1 kHz.

Reference is sometimes made herein to providing techniques to aid in thedetection of osteoporosis. It should be recognized that references madeherein to any specific condition, disease, bone or bone region are madeto provide clarity in the description and should not be construed aslimiting. It should be appreciated that the techniques of the presentinvention can be equally applied to aid in the detection, diagnosis andevaluation of a variety of conditions or diseases other thanosteoporosis including but not limited to osteomalacia (rickets),osteopenia, Paget's disease, osteoarthritis, osteonecrosis, cancer,fracture and any other metabolic, inflammatory, ischemic, traumatic orinfectious disease or condition of bone. The present invention may alsobe applied in the evaluation of normal conditions of bone such as theassessment of bone growth and remodeling, or in the healing or repair ofa fracture or other condition of bone, and in metabolic andendocrinological or genetic disorders or abnormalities of bone.

Reference is also made herein to measurements of bone organic matrixdensity for a particular bone region (e.g. the hip, the femoral neck,the wrist or a vertebral body) and that such measurements aid in thediagnosis and treatment of osteoporosis. It should also be understoodthat the apparatus and techniques of the present invention are notlimited to computation of bone matrix density in any particular type ofbone nor to diagnosis and treatment of osteoporosis. It is recognizedherein that the techniques of the present invention may be applied toany type of bone including but not limited to cortical bone andtrabecular (cancellous) bone and other bone regions and that examinationof bone matrix density in other bone regions may be useful to aid in thediagnosis and treatment of osteoporosis and/or conditions other thanosteoporosis. In addition, the techniques of the present invention maybe used to detect, diagnose or evaluate conditions of other regions ofthe body containing bone-like tissues or materials such as implants,calcified atherosclerotic plaques, calcified heart valves, and tissuescontaining solid or semisolid collagen such as cartilage, tendon,ligament, scar and other fibrotic tissues. Examination of specimens ofsuch tissues or bone outside of the body, and of synthetic materialscontaining fluid and solid phases are also within the scope of theinvention.

Generally, the system and techniques described herein enable measurementof bone matrix density by first suppressing fluid (e.g. water and fatsignals) with a sequence of frequency selective RF pulses which precedethe pulses of a solid state imaging sequence such as the field gradientpulse and RF pulse of a projection imaging pulse sequence. Because thefluid (water and fat) spectral line widths are relatively narrowcompared to the line width of the bone organic matrix, and because theyoccur at different Larmor frequencies (separated by the chemical shiftdifference between water and fat), they may be excited with so-called“weak” RF pulses, which excite only a relatively narrow band offrequencies. In principle, either water or fat could be excited with aweak 90° RF pulse, which would convert all the signal of that substance,for instance water, into transverse magnetization which would then beallowed to dephase under a field gradient pulse (the dephasinggradient). In practical systems, however, it is nearly impossible toapply a perfect 90° pulse to an entire volume of a specimen or subject.Therefore, the pulse sequence may be repeated with the addition of a180° frequency selective pulse following the first dephasing gradient.This inverts any residual water longitudinal magnetization to itsinverse, which then contributes to the image, but in a negative sense.Subsequent repetitions of the pulse sequence are performed alternatelywith and then without the 180° pulse and its accompanying dephasinggradient, and all of the recorded signals are co-added. The residualwater signals, already reduced substantially because of the 90° RF pulseand dephasing gradient pulse, are alternately positive and negative inthe summation but of approximately the same magnitude, and so theylargely cancel. By this method, up to about 97 percent of the watersignal may be suppressed. The same process is applied to suppress thefluid fat signal. In the combined suppression, the first pulse sequencerepetition contains frequency selective 90 pulses for water and fat,accompanied by their gradient dephasing pulses, while the secondrepetition also includes the 180° pulses for water and fat. The pairs ofpulse sequences are co-added to suppress the fluid water and fatsignals. By careful choice of the frequency selective pulse amplitudes(which approximately determine the frequency bandwidth over which thepulses are effective), the effect of these pulses on the signal from thesolid matrix may be reduced while retaining good fluid signalsuppression.

By being able to identify 3-D volumetric bone organic matrix density,the apparatus and techniques of the present invention can be used tocompute a bone organic matrix density and thus aid in the diagnosis anddetection of osteoporosis and a variety of different conditions insubjects. For example, by reliably measuring the bone matrix density ina particular region of interest in a subject, it may be possible toprovide a non-invasive technique for assisting in the diagnoses of avariety of diseases including but not limited to osteoporosis,osteomalacia (rickets), osteopenia, Paget's disease, osteoarthritis,osteonecrosis, cancer, fracture and any other metabolic, inflammatory,ischemic, traumatic or infectious disease or condition of bone. Thepresent invention may also be applied in the evaluation of normalconditions of bone such as the assessment of bone growth and remodeling,or in the healing or repair of a fracture or other metabolic, genetic orendocrinological disturbances of bone.

Turning now to FIG. 2, a magnetic resonance imaging (MRI) system 10 thatmay be programmed to non-invasively measure bone matrix density and aidin the diagnosis and detection of osteoporosis and other conditions inaccordance with the present invention includes a magnet 11 havinggradient coils 12 and RF coils 14 disposed thereabout in a particularmanner to provide a magnet system 15. In response to control signalsprovided from a controller processor 16, a transmitter 17 provides atransmit signal to the RF coil 14 through an RF power amplifier 18. Agradient amplifier 20 provides a signal to the gradient coils 12 also inresponse to signals provided by the control processor 16.

The magnet system 15 is driven by the transmitter 17 and amplifiers 18,20. The transmitter 17 generates a radio frequency drive signal which isamplified by the RF amplifier 18 and applied to the RF coil 14. Thegradient amplifier 20 provides a magnetic field gradient drive signalwhich is applied to the set of gradient coils 12. The resultant magneticfield gradient may have an arbitrary direction. For generating auniform, steady main magnetic field required for MRI, the magnet system11 may be provided by resistive coils driven by a generator, permanentmagnets, superconducting coils, or the earth's magnetic field. Themagnetic fields are generated in an examination or scanning space orregion 19 in which the object to be examined is disposed. For example,if the object is a person or patient to be examined, the person orportion of the person to be examined is disposed in the region 19.

The transmitter/amplifier 17,18 drive the RF coil 14. After the RF pulseis applied to the RF coil 14, spin resonance signals are generated inthe object situated in the examination space 19, which signals aredetected and are applied to a receiver 22. Depending upon the measuringtechnique to be executed, the same RF coil 14 can be used for thetransmitter coil and the receiver coil or use can be made of separatecoils for transmission and reception. The detected resonance signals aresampled and digitized in a digitizer 24. Digitizer 24 converts theanalog signals to a stream of digital bits which represent the measureddata and provides the bit stream to the control processor 16.

The control processor 16 processes the resonance signals measured so asto obtain an image of the excited part of the object. A display 26coupled to the control processor 16 is-provided for the display of thereconstructed image. The display 26 may be provided for example as amonitor, a terminal, such as a CRT or flat panel display.

A user provides scan and display operation commands and parameters tothe control processor 16 through a scan interface 28 and a displayoperation interface 30 each of which provide means for a user tointerface with and control the operating parameters of the MRI system 10in a manner well known to those of ordinary skill in the art.

The control processor 16 also has coupled thereto a pulse sequenceprocessor 31 and an image reconstruction processor 35. The pulsesequence processor 31 includes a fluid suppression processor 32 and asolid imaging processor 34. Also coupled to the control processor is adata store 36. Each of the components depicted in FIG. 1, except for thepulse sequence processor 31 are standard equipment in commerciallyavailable MRI systems. It should be appreciated that the MRI system mustbe capable of implementing the pulse sequence provided by pulse sequenceprocessor 31 and the MRI system must also be capable of acquiring theresultant data.

In some embodiments, the pulse sequence processor 31 may be provided asa general purpose processor or computer programmed to provide pulsesequences in accordance with the techniques described herein. In oneembodiment, the pulse sequence processor 31, fluid suppression processor32 and solid imaging processor 34 may be implemented as computer programcode executed by the same physical processor or by physically separateprocessors. It should also be appreciated that one or more of theprocessors 31, 32, 34 (or all of the processors 31, 32, 34) may beimplemented in hardware (e.g. as an integrated circuit such as anapplication specific integrated circuit (ASIC)) or a combination ofhardware and software.

In some applications it may be desirable to provide a single processoror computer which is appropriately programmed to perform the functionsof control processor 16, pulse sequence processor 31, fluid suppressionprocessor 32 and solid imaging processor 34. In other embodiments,control processor 16, pulse sequence processor 31, fluid suppressionprocessor 32 and solid imaging processor 34 may be provided as speciallydesigned processors (e.g. digital signal processors) or other speciallydesigned digital or analog circuits. In any event, pulse sequenceprocessor 31 is unique in that it is programmed or otherwise designed toprovide a sequence of pulses to allow measurement of the volumetric, 3-Dbone organic matrix density and the detection of osteoporosis and otherdiseases in accordance with the present invention as described below.

The fluid suppression processor 32 and solid imaging processor 34cooperate to provide a sequence of pulses for measuring the volumetric,3-D bone organic matrix density. One particular pulse sequence isdescribed below in conjunction with FIG. 3. Suffice it here to say,however, that in accordance with the present invention, fluid water andfat signals are suppressed with a sequence of frequency selective RFpulses which precede the pulses of a solid state imaging sequence suchas the field gradient pulse and RF pulse of a projection imaging pulsesequence. The fluid (e.g. water and fat) chemical shift selective pulsesare long duration low power rectangular RF pulses. After each selectivepulse, a strong dephasing gradient destroys any transverse water or fatmagnetization. Residual water or fat longitudinal magnetization isinverted on alternate scans, and is cancelled when the scans areco-added. After the train of selective RF and gradient pulses, a shortduration high power nonselective RF pulse excites a portion of theremaining z-magnetization, which should consist of only short T₂*constituents (mostly solid collagen), to be read out as a free inductiondecay (FID) in the presence of a frequency-encoding 3D projectiongradient pulse. The delay between the projection gradient ramp-up andthe hard RF pulse (typically a few hundred microseconds) is chosen topermit any gradient transients to settle. By reading an FID rather thanan echo, and dispensing with slice selection, the pulse sequence isdesigned to image materials with T₂ values far below the minimum echotimes of most scanners.

Referring now to FIG. 3, a flow diagram shows the processing performedby a processing apparatus which may, for example, be provided as part ofan MRI system such as that shown in FIG. 1 to determine the volumetric,3-D bone organic matrix density. The rectangular elements in the flowdiagram are herein denoted “processing blocks” and represent computersoftware instructions or groups of instructions.

Alternatively, the processing blocks represent steps performed byfunctionally equivalent circuits such as a digital signal processorcircuit or an application specific integrated circuit (ASIC). It shouldbe appreciated that some of the steps described in the flow diagram maybe implemented via computer software while others may be implemented ina different manner e.g. via hardware or an empirical procedure. The flowdiagrams do not depict the syntax of any particular programminglanguage. Rather, the flow diagrams illustrate the functionalinformation one of ordinary skill in the art requires to fabricatecircuits or to generate computer software to perform the processingrequired of the particular apparatus. It should be noted that manyroutine program elements, such as initialization of loops and variablesand the use of temporary variables are not shown. It will be appreciatedby those of ordinary skill in the art that unless otherwise indicatedherein, the particular sequence of steps described is illustrative onlyand can be varied without departing from the spirit of the invention.

Turning now to FIG. 3, processing begins in an image acquisition phase38 which includes processing blocks 40 and 42. In processing block 40, afirst pulse sequence fragment having a characteristic selected tosuppress one or more fluid resonance signals is provided in an MRIsystem. In the case of imaging a bone to aid in the detection ofosteoporosis, the first pulse sequence fragment suppresses at least twofluid resonance signals. Although one particular pulse sequence fragmentto suppress fluid is described herein below in conjunction with FIGS. 3and 3A, it is recognized that any technique or pulse sequence whicheffectively suppresses fluid signals may also be used.

Processing next proceeds to processing block 42 in which a second pulsesequence fragment having a characteristic selected such that the pulsesequence fragment images at least solids is provided to the MRI system.Since the first pulse sequence fragment substantially suppresses signalsarising from fluids, when the second pulse sequence fragment is used,only signals generated in response to solids remain.

The second pulse sequence fragment can be provided, for example, as thetype used in conventional solid state MRI which, as is known, detectsboth solid and fluid constituents. Since conventional solid state MRItechniques detect both the solid and fluid constituents, the solidconstituents would normally be obscured by the fluid constituents.However, by first applying the pulse sequence fragment which suppressesfluid resonance signals, it is possible to utilize the conventionalsolid state MRI technique to measure the bone matrix. One example ofsuch a pulse sequence fragment is described in U.S. Pat. No. 6,185,444.

It is also recognized that the second pulse sequence fragment may beprovided as the type which detects only solid constituents. One exampleof such a pulse sequence fragment is described in U.S. Pat. No.5,539,309.

The scanner loops between boxes 40 and 42. After enough data isacquired, processing proceed to step 43 in which the image isreconstructed and then results are displayed as shown in block 44.

Referring now to FIGS. 4A and 4B which when taken together are referredto as FIG. 4 below, a pulse sequence for bone matrix imaging by fat andwater suppressed proton solid state projection MRI includes a firstpulse sequence fragment 50 selected to suppress fluid (e.g. water andfat) signals. The first pulse sequence fragment 50 precedes a secondpulse sequence fragment 52 selected to image at least solids. It shouldbe appreciated that the first pulse sequence fragment 50 may be providedfrom any imaging technique which suppresses fluid signal response.Similarly, the second pulse sequence fragment 52 may be provided usingany solid state imaging technique which images at least solids. Thetechniques described in the aforementioned U.S. Pat. Nos. 6,185,444 and5,539,309 are appropriate for example.

In the particular example shown in FIG. 4, the first pulse sequencefragment 50 includes a first series of RF pulses 54 a and a first seriesof gradient pulses 54 b while the second pulse sequence fragment 50includes a hard RF pulse 74 provided in the presence of a projectiongradient 80 which yields a resulting Free Induction Decay (FID) signal76.

The first series of RF pulses 54 a includes a first pulse 56corresponding to a water-selective 90° pulse 56. Pulse 56 is selected totarget fluids (i.e. pulse 56 covers fluid state resonance widths) andthus is sometimes referred to as a “weak” or “soft” pulse. Thewater-selective 90° pulse 56 is followed by a water-selective 180° pulse58 which in turn is followed by a fat-selective 90° pulse 60, followedby a fat-selective 180° pulse 62. It should be appreciated thatreference is made herein to water and fat selective 90° and 180° pulses.It is recognized that although the target values are exactly 90° and180°, in practice it is always the case that there is a significant(10-30 degrees or more) distribution in flip angles over the volume ofthe subject or specimen, and the actual central value may be far from90° and 180° degrees.

A series of dephasing gradient pulses 64-70 is provided in conjunctionwith the series of water and fat selective pulses 56-62 and together theseries of pulses 54 a, 54 b provide the first pulse sequence fragment 50which corresponds to a fluid suppression pulse sequence fragment.Ideally, after the first pulse sequence fragment 50 has been applied,all fluid signals have been suppressed. Thus, even though the secondpulse sequence fragment 52 detects both fluids and solids, since thefluid signals have been suppressed by the first pulse sequence fragment50, only solid signals are left to detect.

It should be appreciated that fluid signals are suppressed with thesequence of frequency selective RF pulses 54 a which precede the fieldgradient pulse 80 and RF pulse 74 of the projection imaging pulsesequence 52. Because the spectral line widths of fluids are relativelynarrow compared to the line width of the matrix, and because they occurat different Larmor frequencies (separated by the chemical shiftdifference between water and fat), they may be excited with weak RFpulses, which excite only a narrow band of frequencies.

In principle, either water or fat could be excited with a weak 90° RFpulse, which would convert all the signal of that substance, forinstance water, into transverse magnetization which would then beallowed to dephase under a field gradient pulse (the dephasinggradient). In practical systems, however, it is nearly impossible toapply a perfect 90° pulse to the entire volume of the specimen orsubject. Therefore, the pulse sequence may be repeated with the additionof a 180° frequency selective pulse (e.g. pulse 58) following the firstdephasing gradient 64. This inverts any residual water longitudinalmagnetization to its inverse, which then contributes to the image, butin a negative sense (i.e. the 180° pulses 58, 62 compensate for residuein the 90° pulses 56, 60).

Subsequent repetitions of the pulse sequence are performed alternatelywith and then without the 180° pulses 58, 62 (and the accompanyingdephasing gradients 66, 70), and all of the recorded signals areco-added. The residual water signals, already reduced substantiallybecause of the 90° pulse and dephasing pulse, are alternately positiveand negative in the summation but of approximately of the samemagnitude, and so they largely cancel. By this method, up to aboutninety-seven (97) percent of the water signal may be suppressed. Thesame process can then be applied to suppress the fluid fat signal.

In the combined suppression, the first pulse sequence fragment 54 acontains frequency selective 90° for water and fat, accompanied by theirgradient dephasing pulses, while the second repetition also includes the180° pulses for water and fat. The pairs of pulse sequences are co-addedto suppress the fluid water and fat signals. By careful choice of thefrequency selective pulse amplitudes (which approximately determine thefrequency bandwidth over which the pulses are effective), the effect ofthese pulses on the signal from the solid matrix may be minimized whileretaining good fluid signal suppression.

It should be appreciated that in alternate embodiments, it may bedesirable or necessary to provide a sequence in which gradient pulses66, 70 are present for all scans (alternating the 180° pulses 58, 62 onalternate scans as before). Alternatively, either or both gradientpulses 66, 70 may be inverted in sign with respect to gradient pulses64, 68, or may be of different amplitudes or directions with respect togradient pulses 64, 68. Either of these approaches may result inimproved fluid signal suppression.

In one embodiment, the spacing between each of the RF pulses 56-62 (e.g.the spacing between the falling edge of pulse 56 and the leading edge ofpulse 58) is selected so as not to generate spin echoes in the secondpulse sequence fragment 52. Alternatively, the gradient pulses 54 b maybe individually varied in number, sign, magnitude, direction andspacing, and the RF pulses 54 a may be individually varied in number,magnitude, phase and spacing so as to maximize the suppression of fluidsignals and minimize the suppression of solid signals during the imageacquisition pulse sequence fragment 52. In one particular embodiment,the spacing between the falling edge of pulse 56 and the leading edge ofpulse 58 is approximately 1 millisecond (ms) while the spacing betweenthe falling edge of pulse 58 and the leading edge of pulse 60 isapproximately 2 ms.

It should be noted that the transition between the last dephasinggradient pulse 70 in the first pulse sequence fragment 50 and theprojection gradient 80 in the second pulse sequence fragment 52 may beaccomplished via a transition pulse having a shape such as pulse 78 ortransition pulse having a shape such as pulse 78′. Any shape oftransition pulse 78 is within the scope of the invention. It is onlynecessary that gradient pulse is sufficiently stable by the time thesecond pulse sequence fragment 52 starts such that good image quality isachieved.

In one experiment, a pulse sequence to suppress signals contributed fromboth water and fat (long T₂) and observe signals from the solidcomponents of bone matrix (collagen, short T₂) as shown in FIG. 3 wasused. Low power and long duration frequency selective (resonant withwater and fat sequentially) 90° pulses leave short T₂ magnetization(collagen) along the z-axis while rotating the long T₂ (water or fat)magnetization into the transverse plane; the water and fat magnetizationis subsequently dephased by large gradient pulses. To further suppressthe residual water and fat z-magnetization, which remains because of 90pulse imperfections, a 180 pulse was applied following each 90 pulse onalternate scans. The 180 pulses invert only the long-T₂ water and fatmagnetization, leaving short-T₂ magnetization (collagen) unaffected. Alarge amplitude RF pulse, of a duration short compared to the T₂s of thesharper collagen resonances, is applied to rotate the availablez-magnetization into the transverse plane where it is acquired by athree dimensional projection imaging method. If there is any long T₂signal excited by this hard pulse, it is canceled out in two consecutivescans due to the selective 180 degree pulses. Hence the only signalscontributing to the image will be those arising from short T₂components.

Images were taken on a 4.7 T Bruker/GE (Fremont, Calif., USA) CSI OmegaMR system equipped with an Oxford Instruments (Oxford, UK) 4.7 T 33 cmhorizontal bore magnet. Two different probes were used in differentexperiments to accommodate different size specimens. A 2.5 cm insidediameter solenoid coil was used for the water and oil phantom andtrabecular bone specimens. A birdcage coil with inside diameter of 13 cmwas used for larger cortical bone specimens. This coil produced a highlyhomogeneous B₁ field over the specimen volumes (at the expense of lowerfilling factor). The ¹H Larmor frequency was 200.09 MHz. The water andfat selective suppression 90 degree pulses were 2.5-3.0 ms in duration,and the 180 degree pulses were 5.0-6.0 ms in duration. The 90 degree and180 degree pulses were separated by 1 ms, the water and fat suppressionpulse pairs were separated by 2 ms, and the entire suppression sequencewas separated from the beginning of the gradient ramp 78′ by 2 ms (thesechoices are chosen to interfere with water or fat echo formation duringthe matrix signal acquisition). The short hard pulse used to excite thesolid signal was 10 microseconds in duration.

The FID projection data was acquired under fixed gradient magnitudes (60mT/m) in 998 or 1963 directions (field of view FOV 4 or 8 cm,respectively) at a sampling rate of 5 microseconds or 25 microsecondsper complex point. Effectively, about 64 complex points of the FID wereused in the reconstruction. The number of gradient directions (or views)is not a “round” number, but rather is chosen to provide a uniformpattern of coverage in solid angle. The algorithm used to selectprojection directions distributes them on a series of parallel latituderings equally spaced in azimuthal angle; within each ring views arespaced in equal increments of polar angle, with the number of views ineach ring proportional to the sine of the azimuthal angle. This gives apattern uniformly covering the sphere in which views are allotted toequal solid angles. Gradient pulse rise and fall times were 3 ms.

Repetition times TR were 0.5 s and the FIDs were averaged over twoacquisitions with 180 degree receiver phase cycling. A typicalmeasurement time for 998 projections (obtained in a block of 1024acquisitions, because the Omega software requires the number ofacquisitions in a block to be a power of 2) was 18 minutes, and for 1963projections (2048 acquisitions) was 36 minutes. The fractional isotropicresolution of the reconstruction (the linear resolution element dividedby the field of view) is given by sqrt (pi/N), where N is the number ofprojections. The spatial resolution is therefore 2.2 mm for images withFOV=4 cm (N=998) and 3.2 mm for images with FOV=8 cm (N=1963). A smalladditional time was required for separate collection of the weakgradient acquisitions to fill the center of k-space.

Table 1 shows the proton T₁ and T₂ measurements on individual specimensat the ambient temperature of 9° C. inside the magnet bore.

TABLE 1 T₁(ms) T₂(μs) Dry EDTA 419.0 208.5^(a) Decalcified bone 1 DryEDTA 457.0 176.6^(a) Decalcified bone 2 Dry Tendon 1 929.3 131.8^(a) DryTendon 2 1023.9 124.3 Dry Tendon 3 993.0 119.9^(a) Water Fat Water Fats(ms) (ms) (ms) (ms) Bovine Intramedullary fat 1^(b) 982.4 262.6 20.418.6 Bovine Intramedullary fat 2^(b) 895.0 251.6 21.8 16.9 BovineIntramedullary fat 3^(b) 678.0 294.0 30.0 15.7 ^(a)T₂ obtained by 1/(π ×linewidth) ^(b)From cancellous bone tissues

The average T₁s (± standard deviations) of water and lipid in bovineintramedullary fat are 850±160 ms and 270±20 ms, respectively. Theaverage T₂s of water and fat are 24±5 ms and 17±1.5 ms, respectively.These results are consistent with data published in the literature. Theaverage T₁ and T₂ of dry EDTA and HCl decalcified bone are 440±30 ms and190±20 μs, respectively. The T₂ of dry tendon, 125±6 μs, is shorter thanthat of decalcified bone, while the T₁ is longer at 980±50 ms. Theseresults provide the basis for the choice of the water and fatsuppression pulse lengths.

FIGS. 5A and 5B are plots of transverse M_(y) and longitudinal M_(z)components of magnetization M (originally along z) versus t/T₂ followingperfect rectangular RF pulses. FIG. 5A illustrates the γB₁t=π/2 pulse:the magnetization of small t/T₂ (water and fat) constituents is rotatedinto the transverse plane, while the magnetization of large t/T₂constituents (solid collagen and motionally restricted matrix water) ispreserved along the z-direction. FIG. 5B illustrates the γB₁t=π pulse:The initial magnetization of small t/T₂ constituents is inverted, whilethe magnetization of large t/T₂ constituents is preserved along thez-direction. Little transverse magnetization is created irrespective ofthe value of t/T₂.

According to FIGS. 5A and 5B, in the long T₂ regime, if t/T₂˜0.1, morethan 95% of M would be flipped into the transverse plane by the soft 90°pulse. In the short T₂ regime, if t/T₂˜10, only 80% of M would beunaffected by the soft 90° pulse, but if t/T₂˜20, the percentage of Munaffected would increase to 90%. A pulse length of 2.5-3.0 ms wastherefore selected for the soft 90° pulse (suppression pulse). Thechoice of the 180° pulse length is more heavily influenced by the desireto not affect the short T₂ component when inverting the long T₂component. The 180° soft pulse was chosen to be applied at the samepower level but twice the duration of the soft 90° pulse. When appliedin vivo at 37° C., the increased molecular mobility of water and fatwill make the selective pulses work somewhat more effectively than inthe present experiments.

FIG. 6A demonstrates the water and fat suppression performance on aphantom (a 2.5 cm diameter vial containing a layer of corn oil and alayer of water, FIG. 6A). These experiments used the 2.5 cm solenoidcoil. The regular (no suppression) one pulse proton spectrum of thephantom shows the water and fat peaks 3.5 ppm apart (FIG. 6B). Thespectrum, acquired by water and fat suppressed projection imagingwithout the projection gradients, showed that these two peaks weremostly suppressed (FIG. 6C). One-dimensional proton MRI of the phantomby standard FID projection imaging (FIG. 6D) and water and fatsuppressed projection imaging (FIG. 6E) shows that the profile of thephantom is mostly suppressed by water and fat suppressed projectionimaging. The percentage of suppression was calculated by comparingstandard three-dimensional FID projection (FIG. 6F) and water and fatsuppressed projection imaging (FIG. 6G) image intensities over the majorportion of the images (excluding the bright spots in the boundary areawhere RF inhomogeneity near the coil wires causes the method to fail).This calculation shows that more than 97% of signal from both water andfat was suppressed.

Standard proton three-dimensional FID projection imaging (yielding thetotal proton signal) and WASP1 (yielding the matrix signal) wereperformed on an intact diaphyseal cortical bone specimen containing itssolid intramedullary fat mass (which contains water as well as lipid),using the 13 cm diameter birdcage coil with FOV˜8 cm. Because of itsgreater molecular mobility compared to collagen, the intramedullary fatsignal is much brighter than that of cortical bone in the regularprojection image (FIG. 7A). In the water and fat suppressed projectionimages, the intramedullary fat is no longer visible (FIG. 7B).Quantitatively, the intramedullary fat signal is found to be less than3% of its normal value, while the cortical bone signal is reduced to 40%of its normal value. A significant loss of cortical bone signal isexpected because some solid matrix signal loss results from thesuppression pulses, and because cortical bone contains water and a smallamount of lipid which are fully suppressed. This experiment convincinglydemonstrates that what is observed in water and fat suppressedprojection imaging is basically the solid, lipid free, dry organicmatrix of compact cortical diaphyseal bone.

Regular and water and fat suppressed projection images with FOV˜4 cm(FIG. 8) were obtained from cancellous, trabecular bone specimens. Theintensity of the bone signal in the water and fat suppressed projectionimages is assumed to arise from the solid, lipid free, dry organicmatrix constituents of the bone substance of the trabeculae in thecancellous bone tissue. In order to compare MRI data with chemicalanalysis, the outcome of which is an average over the whole specimen,the MRI signal was averaged over all pixels whose signal is at least 10times higher than the average background pixel. The intensity of drybovine tendon was also measured in the same manner using the sameexperimental procedures as a standard for water and fat suppressedprojection imaging. In water and fat suppressed projection imaging ofdry tendon, the signal is reduced to 70% of its normal value. Theintensity ratio of the water and fat suppressed projection image of thevery young trabecular bone divided by the tendon water and fatsuppressed projection image gives the approximate weight percentage ofthe lipid free, dry bone matrix. The weight percentages, calculated fromthe image intensity ratio, for two specimens are 21% and 18%respectively. Gravimetric analyses of identical specimens taken from thesame bones yielded weight percentages of dry organic matrix in wet bonetissue of 17% and 16% respectively. This reasonable agreement indeterminations of organic matrix content by the two different methodsconfirms the general validity of the water and fat suppressed projectionimaging technique. It should be appreciated in FIG. 6 that the displaybrightness of the water and fat suppressed projection images has beenincreased to show the specimens clearly. However, for purposes ofcalculating the matrix density (directly measured as volumetric 3-Ddata), the actual pixel values are used.

All references cited herein are hereby incorporated herein by referencein their entirety.

Having described preferred embodiments of the invention, it will nowbecome apparent to one of ordinary skill in the art that otherembodiments incorporating their concepts may be used.

For example, although the use of this technique for measuring thevolumetric, 3-D bone organic matrix is likely to be important, otherapplications are possible, and are within the scope of the invention.These other applications include any situation in which at least twosubstances are present which differ in their spin-spin relaxation time,and it is desired to produce an MR image of the shorter spin-spinrelaxation time material without interference from the longer spin-spinrelaxation time material. Examples include without limitation biologicaland nonbiological composite materials such as tendon, cartilage,ligament, plaques, fibrotic tissue, calcified tissue, skin, hair, nail,hoof, cuticle, leather, parchment and other hard or soft or fluid animaltissues and their derivatives, wood and other plant tissues and theirderivatives, fibers, foods, agricultural materials, soil, coal,petroleum, tar, oil shale, minerals, rock, fossils, animal and plantremains and other geophysical or petrochemical materials, liquids,gases, chemicals, polymers, rubbers, ceramics, composite materials,sols, gels, colloids, porous materials and liquid crystalline materials,either singly or in combination. Nuclear isotopes other than protons,including without limitation ²H, ¹³C, ¹⁴N, ¹⁵N, ¹⁷O, ¹⁹F, ²³Na and ³¹P,and other substances with magnetic moments, including without limitationelectrons, neutrons, muons, ferromagnetic, antiferromagnetic andferrimagnetic materials, are also within the scope of the invention.

It should be appreciated that Variations of the fluid (e.g. water andfat) suppression pulse parameters, including numbers, substances to besuppressed, numbers of pulses, pulse durations, interpulse durations,pulse flip angles, pulse RF phases, pulse amplitudes and pulse shapes,whether used in combination with projection imaging or other imagingpulse sequences, irrespective of the number of spatial dimensions, areall within the scope of the invention.

It is felt therefore that these embodiments should not be limited todisclosed embodiments, but rather should be limited only by the spiritand scope of the appended claims.

1. A method for producing an image using a magnetic resonance imaging(MRI) system, the steps comprising: a) performing a first preparatorystage of an MR pulse sequence to suppress a transverse magnetization ofa first spin-species having a first T₂ relaxation time; b) performing afirst imaging stage of the MR pulse sequence using projection imaging tosample a first free induction decay signal from a second spin specieshaving a T₂ relaxation time substantially shorter than the first T₂relaxation time and produce a first image data set; c) performing asecond preparatory stage of the MR pulse sequence to suppress thetransverse magnetization of the first spin-species and invert alongitudinal magnetization of a desired spin-species; d) performing asecond imaging stage of the MR pulse sequence using projection imagingto sample a second free induction decay signal from the second spinspecies and produce a second image data set; and e) reconstructing thefirst and second image data sets to produce an MR image including thesecond spin species that is substantially free of the firstspin-species.
 2. The method as recited in claim 1 wherein the first spinspecies include at least one of fat and water and the second spinspecies includes components from bone.
 3. The method as recited in claim1 wherein the first preparatory stage of the MR pulse sequence includesselective 90 degree RF pulses and the second preparatory stage of the MRpulse sequence includes selective 90 degree RF pulses and 180 degree RFpulses.
 4. The method as recited in claim 3 wherein the selective 90degree RF pulses are approximately 2.5 ms to 3.0 ms in duration.
 5. Themethod as recited in claim 4 wherein the 180 degree RF pulses areapproximately 5.0 ms to 6.0 ms in duration.
 6. The method as recited inclaim 3 wherein the first and second imaging stages include nonselective90 degree RF pulses.
 7. The method as recited in claim 6 wherein thenonselective 90 RF pulses are approximately 10 μs in duration.
 8. Themethod as recited in claim 1 wherein the MR image includes informationrelating to a degree of bone mineralization.
 9. The method as recited inclaim 8 further including f) comparing the MR image produced in step e)to a bone mineral density measurement produced using solid state ³¹Pprojection MRI to determine the degree of bone mineralization.
 10. Themethod as recited in claim 1 wherein steps a) to d) are repeated at agiven repetition time to acquire a desired amount of NMR data before thereconstruction of the MR image at step e).
 11. The method as recited inclaim 10 wherein the given repetition time is approximately 0.5 s.
 12. Amethod for producing an image using a magnetic resonance imaging (MRI)system, the steps comprising: a) performing a first preparatory stage ofan MR pulse sequence to suppress a magnetization of a first spin-specieshaving a first T₂ relaxation time; b) performing a first imaging stageof the MR pulse sequence to sample a first free induction decay signalfrom a second spin species having a T₂ relaxation time substantiallyshorter than the first T₂ relaxation time and produce a first image dataset; c) performing a second preparatory stage of the MR pulse sequenceto suppress a magnetization of the first spin-species and invert amagnetization of a desired spin-species; d) performing a second imagingstage of the MR pulse sequence to sample a second free induction decaysignal from the second spin species and produce a second image data set;and e) reconstructing the first and second image data sets to produce anMR image including the second spin species that is substantially free ofthe first spin-species to illustrate a degree of bone mineralization.13. The method as recited in claim 12 wherein the first spin speciesinclude at least one of fat and water and the second spin speciesincludes bone.
 14. The method as recited in claim 12 wherein the firstpreparatory stage of the MR pulse sequence includes selective 90 degreeRF pulses and the second preparatory stage of the MR pulse sequenceincludes selective 90 degree RF pulses and 180 degree RF pulses.
 15. Themethod as recited in claim 14 wherein the first and second imagingstages include nonselective 90 degree RF pulses.
 16. The method asrecited in claim 2 further including f) comparing the MR image producedin step e) to a bone mineral density measurement produced using solidstate ³¹P projection MRI to quantify the degree of bone mineralization.17. The method as recited in claim 12 wherein step a) includesperforming the first preparatory stage of an MR pulse sequence tosuppress a transverse magnetization of the first spin-species having afirst T₂ relaxation time, and step c) includes performing the secondpreparatory stage of the MR pulse sequence to suppress a transversemagnetization of the first spin-species and invert a longitudinalmagnetization of the desired spin-species.